1. Field of the Invention
The present invention relates to radiographic equipment, and particularly, to an X-ray tube having a rotating and linearly translating anode.
2. Description of the Related Art
An X-ray tube is a vacuum tube that produces X-rays, typically found in medical X-ray machines and the like. As with any vacuum tube, there is an emitter, typically a filament cathode, which emits electrons into the vacuum, and an anode to collect the electrons, thus establishing a flow of electrical current, referred to as the “beam”, through the tube. A high voltage power source, for example 30 to 150 kV), is connected across the cathode and anode to accelerate the electrons. The X-ray spectrum produced depends on the anode material and the accelerating voltage.
Electrons from the cathode collide with a target deposited on the anode, with the target often formed from tungsten, molybdenum or copper. During collisions, the electrons lose energy in both collisional and radiative modes. About 1% of the kinetic energy during the collision process is converted into X-ray radiation. This is due to the deceleration of the electrons within the electrical field of the nucleus, or through the creation of vacancies in the inner shells of bound electrons.
FIG. 2 illustrates a typical, prior art Coolidge X-ray tube 100, also referred to as a “hot cathode tube”. The Coolidge tube 100 is a vacuum tube, typically formed from a glass shell 104, having a vacuum formed therein, typically along the order of approximately 10−4 Pa or 10−6 Torr. In the Coolidge tube 100, electrons are produced via the thermionic effect from a tungsten filament 102 heated by an electric current (shown in FIG. 2 as being produced by voltage source VH). The filament 102 forms the cathode of the tube 100. A high voltage potential is produced between the cathode and an anode 106 of the tube (produced in FIG. 2 by high voltage source VC-A), so that the electrons generated by filament 102 are accelerated toward anode 106, and then strike the anode 106 to produce X-rays X. In FIG. 2, the Coolidge tube 100 is shown as also including a cooling device 108, with a water inlet Win and a water outlet Wout, for cooling the anode 106, which heats during X-ray production.
Coolidge tubes are formed as either end-window tubes or side-window tubes. In an end-window tube, the filament is wrapped about the anode, so the electrons have a curved path. The tube 100 of FIG. 2 is a side-window tube. In side-window tubes, an electrostatic lens is used to focus the beam onto a very small spot on the anode 106. The anode 106 is specially designed to dissipate the heat and wear resulting from this intense focused barrage of electrons. The anode is precisely angled at between 1 and 20° off perpendicular to the electron current so as to allow escape of some of the X-ray photons X which are emitted essentially perpendicular to the direction of the electron current. The anode is typically made from tungsten or molybdenum. Further, the tube has a window designed for escape of the generated X-ray photons. The input power of a typical Coolidge tube usually ranges from between 1 and 4 kW. Exemplary Coolidge X-ray tubes are shown in U.S. Pat. Nos. 1,211,092; 1,251,388; 1,917,099; and 1,946,312, each of which is hereby incorporated by reference in its entirety.
FIG. 3 illustrates a typical, prior art rotating anode tube 200. The rotating anode tube is an improvement of the Coolidge tube. Because X-ray production is very inefficient (99% of incident energy is converted to heat), the dissipation of heat at the focal spot of the electron beam is one of the main limitations on the power which can be applied. By sweeping the anode past the focal spot, the heat load can be spread over a larger area, greatly increasing the power rating. With the exception of dental X-ray tubes, almost all medical X-ray tubes are of this type.
The rotating anode tube 200 is also a vacuum tube, formed from shell 202 having an X-ray window 210 formed therein. The anode 204 consists of a disc with an annular target 206 formed thereon. The anode disc 204 is supported on an axle 214, which is supported by bearings 212 within the tube shell 202. The anode 204 can then be rotated by electromagnetic induction from a series of stator windings outside the evacuated tube.
Because the entire anode assembly has to be contained within the evacuated tube shell 202, heat removal is a serious problem, further exacerbated by the higher power rating available. Direct cooling by conduction or convection, as in the Coolidge tube, is difficult. In most tubes, the anode 204 is suspended on ball bearings with silver powder lubrication, which provides almost negligible cooling by conduction.
The anode 204 must be constructed of high temperature materials. The focal spot temperature caused by electrons generated by cathode 208 impinging upon target 206 can reach 2500° C. during an exposure, and the anode assembly can reach 1000° C. following a series of large exposures. Typical materials used to form the anode are a tungsten-rhenium target 206 on a molybdenum core, backed with graphite. The rhenium makes the tungsten more ductile and resistant to wear from impact of the electron beams. The molybdenum conducts heat from the target. The graphite provides thermal storage for the anode, and minimizes the rotating mass of the anode.
Increasing demand for high-performance CT scanning and angiography systems has driven development of very high performance medical X-ray tubes. Contemporary CT tubes have power ratings of up to 100 kW and anode heat capacity of 6 Mj, yet retain an effective focal spot area of less than 1 mm2. Exemplary rotating anode X-ray tubes are shown in U.S. Pat. Nos. 1,192,706; 1,621,926; and 3,646,380, each of which is hereby incorporated by reference in its entirety.
In typical X-ray tubes, such as those described above, approximately 1% of the energy of the electron beam is converted to useful X-ray radiation, with 99% of the energy being lost as thermal energy. Thermal loss is of particular importance in high definition imaging, in which the electron beam must be focused on as small a target area as possible over a time period that is as short as possible. Image resolution depends upon both factors in diagnostic X-ray systems. Thermal energy gain within the target is a serious obstacle to the reduction of electron beam size or shortened exposure time.
Excess heat may be removed via conduction, as described above with reference to Coolidge tube 100, or the problem of instantaneous heating may be at least partially controlled by rotating the anode, as in rotating anode tube 200. Such solutions, however, only offer one degree of freedom in heat spreading. It would be desirable to provide an X-ray tube that can provide two degrees of freedom of heat dissipation, allowing for much higher instantaneous power limits.
Thus, an X-ray tube having a rotating and linearly translating anode solving the aforementioned problems is desired.